This disclosure relates generally to diagnostic imaging and, more particularly, to improved power conversion for a computed tomography (CT) system.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan or cone-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are transmitted to the data processing system for image reconstruction. Imaging data may be obtained using x-rays that are generated at a single polychromatic energy. However, some systems may obtain multi-energy images that provide additional information for generating images, and therefore include fast switching of x-ray tube kV between two discrete levels of, for instance, 80 and 140 kV. And, it is generally desired to have a crisp transition between the high and low frequencies and if the switching is too slow, blurring can occur in reconstructed images.
The x-ray generator of a CT system is typically located within the gantry and, as such, rotates about an imaging bore during data acquisition on a rotatable side of the gantry. The x-ray generator includes the x-ray source, a high voltage (HV) tank, and an inverter that is operationally connected to a slip ring. External to the slip ring and on the stationary side of the gantry is a power distribution unit (PDU). The inverter is typically fed with a DC voltage, for example, 650 VDC, and generates an AC waveform of, for example, approximately 300 VAC, at a frequency of typically 20-50 kHz. The AC frequency is fed to the HV tank, which has a transformer and rectifiers that develop a DC HV potential. The HV potential is applied to the x-ray source.
According to one known configuration, the inverter is positioned on the rotating base and therefore rotates with data acquisition components. The inverter includes, in one known arrangement, a full-bridge or “H” configuration of four (4) power switches that switch in a pattern to control the inverter current, and thus output power of the converter. The full-bridge includes two legs, each of which includes an upper and a lower switch. The switches are typically insulated-gate bipolar transistors or IGBTs. In this known configuration of four power switches, switching in the two legs is controlled in a pattern such that either the upper or lower switch of each leg is on. The switching between on and off is controlled in such a fashion that a high-frequency inverter current is formed, which is in turn fed to the HV tank, as stated. One known switching frequency of a four-switch design is 50 kHz, however as detector scintillator technology and other system operating parameters have increased, so too has the need to operate at a higher frequency.
Power converter performance is generally limited due to its maximum switching frequency, which is a function of the switch technology used. Further, some converters are designed to provide a changing output voltage waveform, the fidelity of which is limited by the switching frequency (e.g., fast kV switching for dual energy imaging). Increasing the switching frequency may allow decreasing magnetic core area needed for a given power. Some known converters use switching devices that includes MOSFETS and silicon carbide MOSFETS, but such devices are relatively very expensive and can be difficult to package.
Therefore, it would be desirable to have an improved resonant converter in a CT system.